This invention relates to catheter-mounted ultrasonic transducers of the phased-array type, and in particular to a method for manufacturing an end portion surrounding a catheter-mounted phased-array ultrasound transducer.
Catheter-mounted ultrasonic transducers have in the past taken several forms, including (1) single-element transducer crystals that are pointed radially outward and rotated about the axis of the catheter, (2) radial array transducers, and (3) linear array transducers. Bom U.S. Pat. No. 3,938,502 discloses one catheter-mounted ultrasonic array which utilizes a radial array arranged circumferentially around the axis of the catheter. Proudian U.S. Pat. No. 4,917,097 describes a similar radial array (and alludes to other geometries) that require multiplexing of the ultrasound signals near the elements of the array. Seward et al. (Seward, J. B., D. L. Packer, R. C. Chan, M. G. Curley, A. J. Tajik (1996), "Ultrasound Cardioscopy: Embarking on a New Journey," Mayo Clinic Proceedings, 71(7)) have described a phased array transducer for insertion into the heart. Such an array has the advantage of increased power: as the transducer array is made longer, the number of elements can be increased, thereby increasing the total radiation area.
Conventional phased-array, linear ultrasonic transducers are typically constructed using a piezoelectric material such as PZT. The piezoelectric material is formed into individual elements, arranged side by side with the lengths of individual elements parallel to one another. PZT is typically built on a backing material that reflects most of the ultrasound energy generated by the PZT, and also tends to absorb energy that is coupled into it. The active surface of the PZT is covered with a second material, called the matching layer, that couples ultrasonic energy from the PZT into the tissue that the transducer is in contact with. The backing material and the matching layer are typically made of composite material such as epoxy loaded with a heavier material such as alumina. By adjusting the phase of waveforms applied to the PZT elements, ultrasonic energy can be focused and steered within a plane oriented parallel to the array and the catheter axis. The techniques for designing transducers and steering them are discussed in texts such as Kino (Acoustic Waves, Prentice Hall, Englewood Cliffs, 1987) and Wells (Biomedical Ultrasonics, Academic Press, London, 1977).
FIG. 7 shows a prior-art, linear, phased-array transducer, and identifies the X, Y and Z coordinates for this transducer. In FIG. 7, the PZT material is identified by the reference symbol P, the backing material by the reference symbol B, and the matching layer by the reference symbol M. By properly controlling phase of the transducer signals applied to the individual piezoelectric elements P, the location and size of the focal spot in the XZ plane can be controlled. The size of the focal spot in the Y dimension is typically determined by a lens applied to the transducer. Such a lens focuses ultrasonic energy in the Y direction by taking advantage of the difference in the speed of sound in the lens material and in tissue in contact with the lens. If a lens has a speed of sound that is slower than that of adjacent tissue and is convex in shape, ultrasonic energy is caused to converge in the ZY plane. The ultrasonic energy focuses in a spot that is spaced from the piezoelectric elements P by a distance controlled by the radius of curvature of the lens and also by the difference in speed of sound between the lens and the adjacent tissue. As the speed of sound of the lens is made increasingly slower than that of adjacent tissue or as the radius of curvature of the lens is made progressively smaller, the focal spot approaches more closely to the transducer. When the focal spot is positioned close to the transducer, the width of the ultrasonic field rapidly diverges as the wave propagates past the focal spot. Of course, if the speed of sound in the lens is faster than in the adjacent tissue, the lens material would be formed in a concave shape to obtain the desired focusing.
It is desirable to maintain the width of the ultrasound field as thin as possible in the Y dimension. This keeps the intensity of the ultrasound energy as high as possible, which increases the strength of the reflected signal when the ultrasound is reflected by structures in the tissue. It is also desirable to keep the thickness of the field in the Y dimension as uniform as possible as the ultrasound propagates in the Z direction. This is because reflections of ultrasound energy at any particular depth that are detected at the transducer represent the integrated reflected energy within the ultrasound wave at that depth. If the ultrasound field is too wide, an object causing a reflection may be indistinguishable from the surrounding tissue. If the thickness of the ultrasound field varies from thin to thick as it propagates in Z, then an object that might be detected where the field is thin might not be detected where the field is wide, which is confusing and counterintuitive to the physician. Thus, it is desirable to maintain a thin, but uniform, ultrasound field width in the Y dimension as the wave propagates in the Z dimension.
Seward, et al. (Seward, J. B., D. L. Packer, R. C. Chan, M. G. Curley, A. J. Tajik (1996), "Ultrasound Cardioscopy: Embarking on a New Journey," Mayo Clinic Proceedings, 71(7)) have described a phased array ultrasound transducer for insertion into the heart. This transducer utilizes a linear phased array of the type shown in FIG. 7, and it offers many improvements over catheter-based radial imaging transducers of the past. These advantages are detailed in the Seward paper, but can be briefly listed as follows: the image plane is advantageous when imaging therapeutic interventions in the heart; the overall aperture of the transducer is large, improving the ultrasound energy and the penetration depth of the tissue; and the transducer is compatible with modern ultrasonic scanning systems.
The Seward transducer is made of conventional materials, including an epoxy-based backing block and a silicone-based lens. The transducer is constructed of 128 elements operating at 5 or 7 MHz. The total array extends for 14 mm in the X direction and 3.3 mm in the Y direction. The backing block is 5 mm in depth or more. As such, the overall diameter of this catheter is 8 mm. If the lens were formed into a cylinder with an 8 mm diameter, it would cause the ultrasound focus to be too close to the transducer, and the ultrasound field would then begin to diverge quickly, causing a loss of image quality and a loss of sensitivity and penetration depth. For this reason, the lens of the Seward transducer is flattened in the region of the transducer, making the forming of the final catheter more difficult.
A need presently exists for a catheter mounted, linear, phased-array transducer that is more easily manufactured.